The present invention relates to ventilator and heat exchange systems and, more particularly, to a xe2x80x9cmixed-modexe2x80x9d gas-plus-liquid ventilator system using an endotracheal catheter to add and remove liquid ventilation or heat-exchange medium from the lungs continuously and/or cyclically, with delivery of gas to the lungs at a rate and volume independent of addition and removal of liquid.
The present invention relates to ventilator and heat exchange systems and, more particularly, to a xe2x80x9cmixed-modexe2x80x9d gas-plus-liquid ventilator system using an endotracheal catheter to add and remove liquid ventilation or heat-exchange medium from the lungs continuously and/or cyclically, with delivery of gas to the lungs at a rate and volume independent of addition and removal of liquid.
There are many situations in both human and veterinary medicine where it is desirable to rapidly induce or reverse hypothermia. There are also many clinical situations where it is essential to be able to rapidly reduce dangerously elevated body temperatures to near normal, as in the case of hyperthermia from heat stroke, drug or surgical anesthetic reaction, and febrile illness secondary to stroke, infection or other illnesses. In fact, it has been demonstrated in a number of studies that patient mortality is directly dependent on high temperature exposure time, and inversely dependent on the rapidity with which core temperature-is normalized.
Heretofore, the only clinically available means of achieving very rapid reduction in body temperature (or conversely, of re-warming from hypothermic temperatures) has been the use of invasive methods of heat exchange, such as cardiopulmonary bypass (circulating blood over a heat exchanger), or peritoneal and/or pleural lavage. A third, slower alternative for changing body temperature involves immersing the patient in a bath of heated or chilled liquid or gas (e.g. helium). The problems with these approaches are many:
1) External means of chilling or re-warming are relatively slow ( less than 0.01xc2x0 C. to 0.20xc2x0 C./min), and produce a host of undesirable and sometimes lethal complications. In the case of cooling, the chilling of external body tissues results in vasoconstriction which interferes with the delivery and removal of oxygen, nutrients, and wastes from the peripheral tissues.
2) Re-warming from hypothermia by external means can cause a peripheral vaso-relaxation, hypotension, and effective hypovolemia, for which the cold-impaired heart and autonomic nervous system cannot compensate. Profound hypotension may develop, causing cardiac arrest and death, sometimes paradoxically in people presenting for apparently non-critical conditions.
3) Peripheral tissues being re-warmed recover the need for oxygen and metabolic substrates before the circulatory system and other organs can deliver them (since these organs are still cold and functioning marginally). This resulting imbalance between metabolic supply and demand results in the generation of large amounts of anaerobic waste products, including carbon dioxide and lactate, which decrease blood and tissue pH and result in severely disturbed homeostasis.
4) If external re-warming proceeds without inducing cardiac arrest, a second phase of risk occurs when xe2x80x9cafter-dropxe2x80x9d is experienced. After-drop is a reduction in body core temperature during slow external re-warming. After-drop occurs as a result of peripheral vasodilation during patient re-warming, thus allowing large amounts of blood to flow through deeply chilled peripheral tissues, resulting in a seemingly paradoxical drop in body core temperature. After-drop can result in cardiac arrest during patient re-warming if the heart is cooled below its critical threshold for fibrillation. Though some controversy exists about the relative importance of this process in humans, it still remains of great concern to specialists in the field.
5) The use of invasive temperature modifying techniques such as peritoneal and pleural lavage, extracorporeal perfusion, or central venous cooling, are either not very effective (e.g. lavage techniques), or can be performed only in a medical setting by highly skilled, licensed practitioners (e.g. physicians). Most importantly, these techniques cannot be safely or reliably performed in the field by paramedics or other non-physician emergency medical personnel. In the case of techniques which require vascular access, many medical facilities possess neither the complex and costly equipment required to carry out such procedures, nor the highly skilled personnel necessary to perform such procedures. A particular problem with these methods is the need for bulky, complicated, failure-prone equipment which may be difficult to store in states of readiness (e.g. cardiopulmonary bypass apparatus). Technical errors and mechanical failures associated with extracorporeal techniques carry a high risk of morbidity, with such errors frequently resulting in neurological damage or loss of life.
An alternative to invasive temperature modifying techniques would be to use the large surface area of the lungs as a heat exchanger. Nearly all of the cardiac output (i.e., all blood flowing to the body) flows through the lungs, and since the lungs possess a surface area of at least 70 square meters, they form an ideal heat exchanger that would allow for rapid core cooling and re-warming of the patient without the problems associated with the techniques previously discussed. In addition, since the lungs are accessible via the trachea, the relatively benign maneuver of endotracheal intubation (a skill universally possessed by paramedics) allows for quick field access to this powerful heat exchanger. The potential utility of the lungs as a heat exchanger was first recognized by Clark and Gollan in the 1960""s, when they used the perfluorochemical FX-80 to demonstrate the concept of total liquid breathing in mice. The concept of using the lungs as a heat exchanger for therapeutic purposes was first proposed by Shaffer et al. in 1984, using total liquid ventilation and the fluorocarbon xe2x80x9cRimar 101xe2x80x9d (Rimar Chimica S.p.A., Vincenza, Italy).
Heat exchange in the lungs using liquid ventilation is superior to gas ventilation because at standard temperature and pressure, gases such as oxygen and air have only approximately a 2200th of the volumetric specific heat capacity of water. Thus, under ordinary circumstances the lungs serve as a relatively poor heat exchanger if only gaseous media are used. This includes the use of the highly conductivexe2x80x94low viscosity gas mixture of oxygen and helium (Heliox). The high conductivity of Heliox makes it far more efficacious as a heat exchange medium under high pressure conditions where its specific heat capacity is greater that at normal pressures; however these conditions are of little relevance to most clinical situations.
Liquid ventilation involves the breathing of gas-carrying liquid as the medium of gas exchange within the lungs. Since the first liquid ventilation experiments (1950""s) in mice using super-oxygenated saline, several liquid media for ventilation have been studied. The class of agents currently optimized to function as liquid breathing media are the fluorocarbons (containing only fluorine and carbon), and the organic perfluorochemicals. (PFCs). PFC compounds contain elements other than fluorine and carbon, with fluorine or other halogens comprising the majority of peripheral moieties. within the molecule. As a class, PFC compounds comprise molecules that are relatively insoluble in either water or lipid, and are more-or-less chemically and pharmacologically inert. PFCs do not dissolve native lung surfactants, and are far less injurious to the lungs than any known silicone or water-based solution.
Historically, the first mode of liquid ventilation studied was total liquid ventilation (TLV). In TLV all of the gas within an animal""s lungs is replaced with liquid, and each breathing cycle (tidal volume) is composed entirely of liquid medium. While this modality holds promise in deep-sea diving research, it has not yet been used in humans.
In heat exchange, each TLV breathing cycle provides a certain volume that can be passed through the lung heat exchanger. As noted, the advantage of TLV (as opposed to gas exchange) for heat exchange, is that liquids such as PFCs have a specific heat capacity several thousand-fold that of gas at normal pressure. Despite this advantage, total liquid ventilation suffers from a number of drawbacks:
1) In TLV it is necessary to completely eliminate air from the animal""s lungs and the ventilating circuit, because PFCs do not pump well in many systems due to xe2x80x9cvaporlockxe2x80x9d. The maneuvers necessary to clear all gas from the system are problematic and time consuming.
2) Due to the increased viscosity of liquids (PFCs are 80 times more viscous than air), the number of liquid breaths attainable per minute is sharply constrained compared to air ventilation. Typically, no more than 5 to 7 liquid breaths per minute are possible (Shaffer T H, et al., 1984). This is approximately one forth the usual rate at which tidal gas volume ventilation occurs in animals of this size.
3) In addition, the maximal liquid ventilatory xe2x80x9cminute volumexe2x80x9d (dV/dt) is more tightly constrained in TLV making adequate gas exchange problematic in situations where oxygen demand is high, and the need to remove CO2 is great. Carbon dioxide removal is a particular problem in TLV because PFCs have a lower carrying capacity for CO2 at physiologic partial pressures (which cannot be changed much), than they have for O2 (which is easily deliverable at artificially high partial pressures).
Miyamoto and Mikami in 1976 calculated that the resting man produces normally 192 mL/min of CO2 (S.T.P.). This level of CO2 production would require TLV (PFC) ventilation volumes of about 4 L/min (or about 70 mL/kg/min). Although this is only 70% of the normal gas ventilatory flow for a resting adult, it is near the upper limit of flows that can be accomplished at normal pressures in TLV (Kylstra, 1974). The higher peak and mean ventilating pressures necessary to move the amount of liquid required for CO2 exchange in TLV would expose the lungs to an increased risk of barotrauma (pressure injury) and volu-trauma (over-distention injury).
During higher than normal CO2 production rates (e.g. disease), TLV would clearly not be adequate for CO2 removal. Examples of high CO2 (hypercapnic) states are 1) increased metabolic states (e.g. cancer, infection, burns), 2) states of physiologic stress (e.g. hyperthermia, agitation), and 3) post-ischemic conditions where substantial metabolic debt has been incurred and the need to rapidly unload CO2 and deliver large amounts of O2 are essential. Such hypercapnic/hypercarbic states are also frequently present in shock due to sepsis or trauma, and thought to be due to both an increased production of CO2, and a decreased elimination of CO2 due to low blood flow or pulmonary edema.
In anesthetized, paralyzed, normothermic dogs, TLV is capable of maintaining steady-state gas exchange with adequate O2 delivery and CO2 removal. However, TLV is not adequate to steady-state CO2 removal under basal metabolic conditions in smaller animals with higher specific metabolic rates, such as guinea pigs. As Matthews and co-workers document (1978), the parameters for maintaining normocapnia in anesthetized beagles are narrow, even under basal normothermic metabolic conditions. In this study, as liquid ventilation rates were increased from 2.8 to 5.6 liquid breaths per minute, and alveolar ventilation was increased from 574 to 600 mL/min/animal (increase of 4%), the paCO2 continued to increase until dangerous hypercapnia occurred. The authors suggested that this increase was due to a 2% drop in liquid-alveolar ventilation, however using their own formulas and data, we have calculated that the dogs receiving higher ventilation rates actually have higher rates of alveolar ventilation (dVa/dt). These results would seem paradoxical until consideration is given to the inverse relationship of paCO2 to alveolar ventilation, a relationship which holds only if equilibrium between blood and xe2x80x9calveolarxe2x80x9d CO2 (actually, alveolar and small airway CO2) is reached for each breath. The fact that high TLV ventilatory rates resulted in rising paCO2 in this paper despite ,increased xe2x80x9calveolarxe2x80x9d ventilation (dVa/dt), indicates that the high ventilatory rate used was too rapid for blood/airway CO2 equilibrium to be reached. This behavior is a limitation of the diffusion speed of CO2 in PFC in the small airways, conduits that are constructed on a size scale for xe2x80x9cgas-in-gasxe2x80x9d diffusion, but not xe2x80x9cgas-in-liquidxe2x80x9d diffusion.
Diffusion limitations on CO2 removal in TLV models have been noted by several research groups. This diffusion limited failure of CO2 equilibrium acts to increase xe2x80x9cdiffusion dead spacexe2x80x9d and in practice places even more constrictive limits on usable ventilation rates with perfluorocarbon (PFC) liquids in TLV. These limits are in addition to those already imposed by the viscosity of the fluid itself. For example, Koen and Shaffer found that TLV in young cats showed maximal CO2 elimination at a ventilatory rate of 3 to 3.5 breaths per minute. Decreasing CO2 clearance occurred at lower rates, due to insufficient ventilation, whereas decreased CO2 clearance occurred at higher rates due to CO2 diffusion limitations with short liquid dwell times. In summary, at higher TLV ventilatory rates, equilibrium in pCO2 between alveolar blood and freshly inspired liquid in the small airways is not reached. For this reason, liquid alveolar ventilation in TLV cannot be arbirarily increased, for fundamental reasons involving pressures limits on high liquid flows, and also diminishing gas exchange at rapid liquid flow rates.
4) PFC viscosity (pressure/flow) also places a limit on the rate at which heat can be extracted from an animal or patient using TLV. In addition to the CO2 diffusion limitation, there is indirect evidence suggesting that thermal equilibrium is not reached between blood and liquid in small airways at high TLV xe2x80x9calveolar ventilationxe2x80x9d rates. Thus, there appears to be a heat-diffusion limitation to TLV that is analogous to the CO2 diffusion limitation.
This phenomenon may explain why Shaffer""s TLV cat studies failed to achieve concomitant increases in the rates of animal core cooling, when significantly greater PFC temperature gradients were used (Shaffer T H, et al., 1984). In Shaffer""s report, it was found that decreasing PFC infusion temperature from approximately 20xc2x0 C. to about 10xc2x0 C. (from xcex94T=15xc2x0 C. to xcex94T=24xc2x0 C.), resulted in cooling rates increasing from 0.13xc2x0 C./min (7.8xc2x0 C./hr) to 0.15xc2x0 C./min (9.0xc2x0 C./hr), a change of only 15%. This 15% increase occurred despite an increase of xcex94T equal to 60%. These results suggest a sharp decline in the efficiency of heat extraction with increased xcex94T at higher TLV ventilation rates (in this experiment, rate was increased from 4.5 to 5.3 liquid breaths/min).
In Shaffer""s study, the authors calculate from PFC inspiration and expiration temperature differences, a 96% increase in heat extraction per kg from their animals at the 10xc2x0 C. PFC infusion temperature versus that calculated at 20xc2x0 C. However, the fact is that this increase in heat extraction does not show up in the rate of body core cooling (15%), to which it should be proportional. This indicates that Shaffer""s calculations of heat removal performed on the basis of integrated measurements of expired fluid temperatures must have been in error. As further evidence of this error, calculations of expected cooling rates of animals used in this study (using a reasonable 0.8 cal/g/xc2x0 C., or kcal/kg/xc2x0 C. average specific heat capacity for the body), indicate that up to half of the heat extraction calculated by PFC temperature differences in this experiment are unaccounted for even at the fastest cooling rates. For example, an animal with an average 0.8 kcal/kg/xc2x0 C. specific heat capacity, cooling at the reported rate of 9.0xc2x0 C./hr, could theoretically give up heat at a rate no faster than (0.8 kcal/kg/C)(4184 J/kcal)(9 C/hr)=30,124 J/kg/hr. However, Shaffer""s experiment reports on the basis of temperature readings of PFC infused and expired, the extraction of 65,637 J/kg/hr. It is likely that the difficult integration of [expired fluid temperature] versus [fluid volume] curve for this experiment was in error by a factor of 2.0. For examples of experiments in which integrated cooling rates calculated from PFC temperature differences match actual animal body cooling, see the canine experiments using hand controlled infusion below. The authors of the present patent have found that at rapid (machine-controlled) liquid infusion and removal rates, peak fluid temperatures do not accurately reflect volume-averaged fluid temperatures, or fluid heats.
The second mode of liquid ventilation to be studied was Partial Liquid Ventilation (PLV). In PLV, the subject""s lungs are partially (usually to functional residual capacity (FRC) of 30 mL/kg of body weight or about ⅓rd of total lung capacity (assumed hereafter to be 90 mL/kg) loaded with PFC liquid. In PLV, PFC liquid loading is accompanied by conventional mechanical ventilation using a standard gas ventilator at normal gas rates and tidal volumes. Since the breaths are delivered as gas, PLV allows for the number of breaths per minute, and alveolar ventilation rates, to be set much closer to the physiologically acceptable and desirable rate. PLV can even be used with high frequency gas ventilators, and can accommodate a wide range of metabolic states in which the demand for O2 delivery and CO2 removal is greater than that of basal states. PLV is currently being tested in human clinical trials.
During PLV, gas exchange occurs across material boundaries at two locations: 1) between the PFC liquid and the circulating blood (across the alveolar membrane), and 2) between the PFC liquid and the ventilating gas in the airways, where a short-lived turbulent foam of PFC and ventilating gas is created. The low viscosity of this PFC foam allows it to reach briefly into the small airways of the bronchial tree with each breath and helps explain the complex and poorly understood mechanism of PLV gas exchange. The turbulent mixing of PFC foam may also explain the newly appreciated heat exchange properties of this modality (as well as those of Mixed-Mode Liquid Ventilation, discussed next). The mixing of air and gas in small airways, which will be discussed more fully later, appears to be key to improving the heat transfer limitations of TLV. We believe that the mixing of PFC and gas disrupts laminar liquid (PFC) flow in small airways by introducing turbulence to the fluid, thereby improving the small-scale (small airway) convection necessary for maximal heat transfer rates.
We introduce in this document the novel application of the unique mixing features of PLV to assist in core body-heat transfer. Specifically, the primary clinical utility of PLV has heretofore been in the treatment of adult and neonatal respiratory distress syndromes. In these pathological conditions, PFC (among other salutary effects) moves down ward in the bronchial tree due to its high density (1.8 to 2.0 times that of water) opening alveoli which are closed as the result of pulmonary edema (fluid in the dependent portions of the lungs).
Significant heat transfer has not been documented using standard PLV because PFC has been historically loaded slowly into the lungs, and once in place, has not been retrieved (cycled). Using a single dose of ⅔rds or more of total lung capacity (TLC) of cold (slightly below 0xc2x0 C.) PFC (60 mL/kg), it is possible in the dog to achieve a uniform core cooling of approximately 1.5xc2x0 C. with only modest injury from baro- and volu-trauma (see Example I data in Part II below). Further cooling of the test subject does not occur unless a new load of cold PFC is instilled into the lungs. Thus, the use of PLV and single PFC loads, even to extreme volumes (i.e., those approaching TLC) is not a viable means for achieving even moderate, controllable, or lasting hypothermia.
Single loads of PFC are not sufficient to induce significant or lasting hypothermia. This problem is compounded more so when contemplating PLV for hypothermic subject re-warming. This is because there are two unavoidable limitations on how warm the delivered liquid can be. The first is that an absolute temperature limit of 42xc2x0 C. exists beyond which hemolysis and acute thermal injury to tissue occurs. The second limit involves the temperature gradient (xcex94T) between the blood and the PFC liquid, which if significantly greater than 5xc2x0 C., exposes the subject to the risk of gas bubble emboli. This risk occurs because the solubility of nitrogen and other gases in plasma is greater at cold temperatures. Upon re-warming chilled blood, the nitrogen (and other gases) come out of solution, forming gas bubbles which can then embolize both the arterial and venous circulation. This same phenomenon occurs in nitrogen saturated. tissues warmed at very rapid rates, and forms the pathological equivalent of the xe2x80x9cbendsxe2x80x9d experienced by deep sea divers breathing nitrogen-containing gas as they decompress too rapidly.
These temperature gradient constraints on warm liquid delivery to hypothermic subjects sharply limits the maximal therapeutic rate of heat transfer achievable by PLV used for re-warming. For example, it is not possible to deliver liquid that is 35xc2x0 C. above body temperature in a subject still warm enough to have a beating heart (typically 25xc2x0 C. or above). In contrast, it is possible (see Part II) to safely deliver liquids that are 35xc2x0 C. cooler than body temperature. Thus, the xcex94T when re-warming from modest but life-threatening hypothermia (i.e., 27xc2x0 C.) is less than 30% of that which can be achieved during the therapeutic induction of hypothermia.
One aspect of the present invention is a mixed-mode liquid ventilation (MMLV) method for gas and/or heat exchange in the lungs (human clinical and veterinary applications). The MMLV method allows mixing of gas and liquid in the small airways of the lungs, producing small-scale liquid mixing in a convection-like process, rapid return of fluid from the lung periphery, and more rapid and efficient transfer of heat, and dissolved gasses, during the practice of ventilation with liquids.
In one embodiment, nitric oxide or nitric oxide donors are administered to facilitate gas and heat exchange.
In another embodiment the gas is helium.
In a further embodiment the liquid ventilation medium is a perfluorocarbon or perfluorochemical.
A further aspect of the present invention is a method of inducing small-scale mixing of liquid heat exchange ventilation media, using other known ventilation methodologies, alone or in combination. These specifically include known types of gas ventilation, including high frequency oscillating ventilation.
Another aspect of the invention is the use of MMLV to treat hypothermic pathologies by heating said liquid ventilation medium, and thus increasing body temperature.
A further aspect of the invention is the use of MMLV to induce hypothermia for medical purposes, or to treat hyperthermic pathologies, by cooling said liquid ventilation medium, and thus decreasing body temperature. In this embodiment, liquid ventilation media may be infused at temperatures as low as xe2x88x9210 C. This is made possible by the presence of thermally buffering PFC already in the lung, as well as the fact that PFC may be a minor volumetric component of the ventilation mix in MMLV, and small infusions are warmed to temperatures above 0 C before freezing or chilling damage can be done to tissues.
Another aspect of the invention is a method of preserving biological material, for example beating-heart cadaveric preparations, using the rapid cooling available with MMLV.
A further aspect of the invention is a method for increasing the efficiency of CPR using MMLV.
A final aspect of the invention is an automatic apparatus for MMLV, which connects to the lungs via the bronchi and uses a computer to control loading and unloading of said oxygenated liquid and gas so that mixing occurs; and also makes use of the computer to insure that pressure limits are not exceeded, and gas ventilation proceeds in a way which most rapidly induces removal of fluid heat exchange media. In this embodiment, the computer controls liquid infusion in such a way as to maximize time integrated arterial/venous temperature differences, for best cooling rates, subject to ventilatory constraints.
In a typical embodiment said computer controls the liquid ventilatory volume delivered and removed (dV/dt) to be about 10% to 50% that typically necessary for gas ventilation.
In a further embodiment the apparatus has a cold reservoir where pre-cooled PFC may be stored.
In another embodiment the apparatus has a heat exchanger for PFC.
In another embodiment the apparatus has an active liquid ventilation system, which may employ a canula able to remove liquid at the same time gas breaths are delivered. Gas and liquid are typically infused and removed through separate concentric tubes in many of the most efficient implimentations of the invention, but may be removed through the same tube.